Cell death often refers to a situation where the cell is unable to maintain homeostasis from the environment. A tissue is an aggregation of morphologically similar cells and associated intercellular matter, and experiences death if the cells forming the tissue die. There are two distinct types of prelethal phases: oncosis and apoptosis. Oncosis of cells usually follows a variety of injuries, such as toxins and hyperthermia. Many of these injuries destroy adenosine triphosphate (hereinafter “ATP”) synthesis and plasma membrane of cells. The lack of ATP and/or cell membrane integrity leads to imbalance in ion, e.g., Na+, influx and efflux, and hence cell death occurs. Apoptosis of cells, on the other hand, is a fundamental cellular process involving the programmed cell death in response to diverse signals such as limb and neural development, neurodegenerative diseases and during the cellular response to radiotherapy and chemotherapy. During apoptosis, nicotinamide adenine dinucleotide phosphate (hereinafter “NAD(P)H”) levels drop as mitochondrial oxidation becomes uncoupled, cytochrome c is released, caspases are activated, deoxyribonucleic acid (hereinafter “DNA”) fragmented and programmed cell death occurs.
In vitro, oncosis may be quantified by various histological techniques such as Live/Dead assay and NAD(P)H staining. In addition to requiring cell or tissue samples from the host, these histological procedures usually are complex and very time-consuming. In vitro, apoptosis may be quantified by a series of static changes in enzyme and biochemical markers, such as cytochrome c release, poly adenosine-diphosphate ribose polymerase (hereinafter “PARP”) cleavage, and DNA fragmentation. These changes can be monitored in terms of a series of stains and/or reactions fixing the cells, by Apotag™, Hoescht 33258 stain, propidium iodide staining, and fluorescence activated cell sorter (hereinafter “FACS”) analysis. Therefore, real time monitoring of the reaction cannot be performed, as the cell or tissue must be killed during the assay.
For in vivo animal testing of novel therapeutics for conditions such as neurodegenerative conditions and cancer, results are mostly based upon sacrificing the animals and performing necropsy and analysis via immunohistochemistry and other pathologic techniques. In certain cases, animal magnetic resonance imaging (hereinafter “MRI”) may allow imaging of the lesion or structures involved with the disease process being studied. All of these techniques are costly, time consuming and labor intensive, and require the sacrifice of large numbers of animals in order to achieve meaningful results.
There are no current techniques available for real time monitoring of the cellular processes leading to cell death in the petri dish (in vitro) or in animal and human therapeutics trials (in vivo).
Optical spectroscopy, such as fluorescence and diffused reflectance spectra, has been shown capable of detecting subtle changes in both biochemical compositions and morphological characteristics of tissues associated with the progression of disease in near real time, where these differences can be used to detect tissue abnormalities and ultimately lead to optical tissue discrimination. Optical spectroscopy has been successfully applied to detect disorders of various organs such as cervix and skin both in vitro and in vivo. Several commercial systems are currently available for clinical diagnosis in the bronchus, cervix, etc. However, relatively few studies have addressed the diagnostic potential of optical spectroscopy in tumors, such as brain tumor and liver tumor. Particularly, optical spectroscopic characteristics of liver tissues have not yet been developed.
Cancers, such as liver cancers, pose a significant problem to public health worldwide. Selecting the best treatment for liver cancer depends on the physician being able to precisely identify the type, location, size and borders of the tumor or tumors. By matching that information to a variety of treatment possibilities and considering the benefits and limitations of each, the physician can select the best course of action. Surgical removal of liver cancer tumors is considered to be the most effective treatment for liver cancer. Unfortunately, about 70% of patients cannot have this surgery due to the size or location of the tumors or other health factors. Thermotherapies such as radio frequency ablation (hereinafter “RFA”) and laser-induced thermotherapy are often considered as alternative treatments. Currently, all thermotherapy procedures suffer from the lack of an adequate feedback control system, making it difficult to know precisely when to cease coagulation.
Proper deployment of an ablation probe is the first requirement for achieving a successful and effective liver tumor ablation (i.e., using the minimum numbers of ablation to achieve a thermal coagulation zone larger than the tumor). Currently, ablation probe placement is approximately determined in accordance with palpation and freehand ultrasound imaging. The accuracy of this approach, unfortunately, is hindered by the limitations of tumor margins detection using tumor echogenicity and stiffness. Once the ablation probe is in place, clinicians usually conduct the ablation using a pre-determined power and therapeutic duration (i.e., heating time) and/or local temperature information to conduct thermotherapies for liver tumors. This practice often yields unsatisfying therapeutic outcomes due to the fact that tissue characteristics vary drastically from patient to patient. In addition, the dynamics of tissue characteristics, such as temperature-dependent tissue properties, further complicate the process of thermal coagulation of liver tissues. Since the coagulation zones are often not visible during thermotherapy, it is difficult, if not impossible, to precisely determine the endpoint of therapy. To avoid the undesired therapeutic consequences and to perform precise and effective ablation, an effective feedback control strategy is clearly needed. This feedback mechanism should provide an objective method of determining the ideal placement of the ablation probe, and an objective endpoint for thermotherapies of liver tumors.
Since thermal coagulation of tissues is an outcome of the interaction between heat and tissue components, it is obvious that local temperature can serve as a convenient metric for monitoring the progress of thermotherapies of liver tumors. This concept has been previously implemented using thermocouple measurements and intraoperative MRI (hereinafter “iMRI”). The translation of local temperature-time history into degree of local tissue thermal damage often requires the assistance of a thermal damage model based on rate process theory. In this model, tissue thermal damage is expressed as an Arrhenius integral that depends on tissue temperature-time history, a tissue-dependent frequency factor (A), and an activation energy barrier (Ea). As indicated by the model, the effectiveness of temperature-based tissue thermal damage assessment hinges on the accurate measurement of local temperature-time history, which is difficult to achieve with either thermocouples or iMRI. In addition, knowledge of A and Ea of tissues is largely unavailable. These limitations make temperature-based feedback control of thermotherapies of liver tumors less than optimal.
Therefore, a heretofore unaddressed need exists in the art to address the aforementioned deficiencies and inadequacies.